Analytical instruments such as biosensors are well established as a means of recording the progress of biomolecular interactions in real time. Biosensors are analytical instruments that employ a variety of transduction technologies in order to detect interactions between biomolecules. Such instrumentation requires microfluidic channels in order to deliver samples to a sensing region. Pumps and valves are preferred to provide a means of moving sample through the channels in a controlled reproducible manner.
Recent interest in microfluidics technology has come about because of a growing need for sophisticated control of fluid streams for such sensing applications. A number of prior systems, referred to as integrated microfluidic cards, are composed of a series of substantially planar substrates possessing channels and structures that when bonded form internal passages and active components such as valves and pumps. Despite much progress these systems are rarely as robust as conventional flow injection analysis fluidic systems where the active components are not integrated into the fluidic card; however, these non-integrated systems typically have relatively large dead volumes.
There are several transducers capable of recording the progress of biomolecular interactions, for example a quartz crystal microbalance. Binding of molecules to the surface of a quartz crystal changes the fundamental resonance frequency which allows quantification of the binding event. Other technologies include light scattering, reflectometric interference spectroscopy, ellipsometry, fluorescence spectroscopy, calorimetry, and evanescent field based optical detection. A particularly effective evanescent field based technology, known as surface plasmon resonance (SPR), exploits the behavior of light upon reflection from a gold-coated optical substrate, for example.
SPR is an optical technique that enables real-time monitoring of changes in the refractive index of a thin film close to the sensing surface where materials to be tested are located (typical material types include a ligand attached to the sensing surface, a fluid buffer, and a fluid analyte which is contained (e.g. soluble or insoluble colloidal solution) in a running/flowing material that is to bind with the ligand and be tested). The evanescent field created at the surface decays exponentially from the surface and falls to one third of its maximum intensity at approximately 300 nanometers (nm) from the surface. Hence the SPR technique is sensitive to surface refractive index changes.
An integrally-formed miniature SPR transducer has been described in U.S. Pat. No. 5,912,456. In this device a photodiode array (PDA) simply records the intensity of the reflected light, from a light emitting diode (LED) over a range of angles. Refractive index changes within the penetration depth of the evanescent field give rise to corresponding changes in the position of the SPR reflectance minimum incident on pixels of the PDA. This change in resonance angle is followed by tracking the change in the pixel position of the reflectance minimum. A minimum tracking algorithm is employed to continuously monitor the position of this minimum as it traverses the photodiode array and the pixel position is then related to a refractive index value. A current configuration of this device possesses three SPR active sensing regions per sensor enabling multichannel operation with real-time reference subtraction. Alternative configurations can allow even more numerous SPR sensing regions.
The delivery of samples to the SPR active sensing regions is made possible by creating flow channels that cover the active sensing regions. Each flow channel possesses an inlet and outlet to allow for the flow of buffer, or samples, over the SPR active sensing regions. The thin film sensing surface is derivatized to possess a polymeric coating that enables biomolecules (“ligands”) to be permanently immobilized on the coating. The immobilized biomolecules usually possess binding specificity for another biomolecule contained in the sample (the “analyte”). The strength of this binding is given by the affinity constant which is simply the ratio of the binding rate constant divided by the dissociation rate constant. It is possible to measure these constants because an SPR-based biosensor records the progress of binding and dissociation events in real time.
Scaling down biochemical analysis instruments has important advantages (for example, sample volume reduction, microfluidics for high mass transport, and mass manufacturable). As a particular example, U.S. Pat. No. 5,376,252 discloses planar microfluidic structures useful for capillary electrophoresis. The channels of these microfluidic structures may have diameters in the range of 50-250 micrometers (μm) and may be manufactured by molding trenches into the surface of a first thermoplastic base and then clamping a second base to the first base which thereby covers the trenches to form channels.
Similarly, U.S. Pat. No. 6,375,871 discloses a method of manufacturing microfluidic devices by attaching two layers with a molded channel pattern in one layer.
An integrally-formed miniature SPR transducer as described in U.S. Pat. No. 5,912,456 is a disposable element. Such a sensor includes a flow channel block that forms a reversible leak tight seal with the sensor's SPR active sensing regions. Permanent attachment of a flow channel to the SPR active sensing region is possible but this can be difficult to achieve without damaging the surface chemistry attached to the active sensing region. Damage can occur due to chemical and/or mechanical damage during the flow channel attachment process. In addition, making reliable, and reversible, fluid connections with the flow channel inputs is difficult.
There are many factors that can have a significant influence on the performance of a biosensor and the quality of the analytical data recorded. Of particular importance are the flow channel's physical properties. Indeed, for an SPR-type sensor the SPR signal is averaged over the SPR active area, and the transport of the biomolecule of interest (analyte) contained in the sample to the SPR active surface results from convective and diffusion forces. These phenomena are described by the mass transfer coefficient (km) and are related to the flow channel dimensions and operational flow rate according to the following expression:
                              k          m                =                  C          ⁢                                                                      D                  2                                ⁢                F                                                              h                  2                                ⁢                wl                                      3                                              (        1        )            where                D=Diffusion coefficient of the analyte (m2/s)        h, w, l=Height, width, and length of flow channel (m)        Bulk flow rate (μl/min)        C=Constant        
Typical kinetic analysis of biomolecular interactions requires that a stable analyte concentration gradient exists and this requires laminar flow conditions. Flow channels with heights that exceed 0.5 millimeter (mm) (500 μm) are often characterized by non-laminar flow conditions. Turbulent flow conditions must be avoided if analytical reproducibility is required. It is apparent from equation (1) that mass transport rates are greatly influenced by the size of the flow channel (particularly the flow channel height) and this has a large effect on the magnitude of binding signals that may be detected. Thus effort should be made to ensure that the flow channel dimensions are minimized. In addition, miniaturization of the flow channel dimensions will ensure that mass transport rates are high. If mass transport rates are low, as is the case with large flow channels, then medium to fast kinetic interaction data will represent mass transport rates and not the kinetic constants for the interaction. This is particularly true when the binding rate of the analyte to the immobilized biomolecule (i.e., ligand) is high.
There is a practical limit to the miniaturization of the flow channel dimensions and this is dictated by considerations of contact area, backpressure, and fluid dynamics, as follows.
Flow Channel Contact Area:
The smaller the flow channel area in contact with the sensing surface, the greater the binding response recorded. If the flow channel area in contact with the sensing surface increases, then the ligand binding response will decrease. This is intuitive as the same number of bound molecules are averaged out over a greater area. The SPR signal records the averaged mass increase per unit area.
Backpressure:
The theoretical flow resistance, R, in a rectangular channel, with a high aspect ratio (i.e., the width is far greater than the height), can be estimated from the Poiseuille slot flow equation where,
                    R        =                              12            ⁢                                                  ⁢            μL                                wh            3                                              (        2        )            where                μ=Solution viscosity        L=Length of channel (i.e., flow channel)        w=Width of channel        h=Height of channel        
The backpressure scales inversely according to the cubic power of the channel height. This shows that decreasing the flow channel height by 2-fold will give an 8-fold increase in resistance to flow. If the resistance is high, then the flow channel seal must be leak-proof above that pressure. Therefore, flow channel heights below 15 μm are not practical for low pressure systems. In addition, flow channels below this height are easily blocked by particulates that are often present in unfiltered samples and buffers. In addition, the internal diameter of tubing used to deliver sample to each flow channel should be greater than 50 μm in order to avoid excess backpressure as backpressure scales inversely according to the sixth power of the radius in tubular channels. In addition excess backpressure causes delays in reaching a steady-state flow. Such delays have significant effects on the performance of the biosensor system. In particular, the time required for a full exchange of running buffer with sample within each SPR flow channel will be delayed. If a bulk refractive index variation exists between the running buffer and sample, then the response due to this bulk index response is difficult to resolve from the actual binding signal, without using a reference channel, unless the exchange time can be reduced (i.e., to less than 10 seconds).
General Fluidic Dynamics Design Principles and Gradients:
A vertical gradient from flow through a vertically oriented small channel is described by a parabolic flow profile where the velocity of the liquid at the walls is zero and the velocity is maximal towards the center of the channel. These velocity gradients cause uneven distribution of analyte binding at the surface. In particular it is important that the SPR active surface is not positioned near the wall of the flow channel where the velocity, and hence analyte binding, is very low. Turbulent flow will exist at the inlet and outlet causing unpredictable analyte binding. Therefore, the SPR active region should be centered along the middle of the flow channel, thereby separating the active sensing regions from the walls, and the inlet/outlet holes.
Also, due to the transit/displacement time through the sensing region of the flow cell, an analyte binding gradient might be caused to exist when a sample is injected into a flow cell. As sample flows into the sensing region, the sample displaces the liquid that is already there. Typically this is a displacement in the direction of flow. Depending on the volume of this region of flow and the flow rate, this displacement may require significant time to carry analyte in the sample across the sensing region. This will cause a gradient in analyte binding to exist along the sensing region in the direction of flow. This gradient can be difficult to account for when processing the data obtained from the flow cell. In some cases it is assumed that this gradient does not exist, but this assumption is only true if the time for complete displacement of the liquid in the flow cell is sufficiently short (e.g., less than one second). For example, if the flow cell volume is 50 nanoliters (nL) and the sample is injected at 50 microliters (μL)/minute, then the flow cell volume is displaced 1,000 times per minute or about 16.6 times per second. The analyte gradient along the flow cell may be neglected in this example. However, if the flow cell volume is 50 μL, then at a flow rate of 50 μL/minute the flow cell is displaced only once each minute, which cannot be neglected.
U.S. Pat. No. 5,376,252 and U.S. Pat. No. 5,313,264 disclose fabricating miniature valves that may be completely integrated into a microfluidic path thereby decreasing prechannel dead volume, which is the volume between the injection point and the flow cell sensing region. This type of dead volume influences the rise and fall times of responses to the sample when it is injected into the flow cell. Accordingly, it is preferable to minimize prechannel dead volume.
Another type of dead volume, interchannel dead volume, typically exists between two sensing channels. This type of dead volume can cause distortions when performing reference curve subtraction and so interchannel dead volume preferably should be minimized. Dead volume between two sensing channels affects the arrival time of sample at the second flow channel. Ideally sample should contact both sensing channels simultaneously if accurate reference curve subtraction is required. A time delay between channels may be accounted for during processing; however, the additional dispersion that occurs in this interchannel dead volume can give rise to response variations between channels if any bulk refractive index difference exists between the sample and the running buffer. In practice such bulk refractive index differences can be expected and reference curve subtraction will therefore introduce artifacts. Thus, such delays decrease the reproducibility of binding responses from one SPR channel to the next and also make reference subtraction more difficult. So the reduction of dead volumes is critical in ensuring high analytical performance. For example, it is usual to use the data from a reference channel to correct data from a working channel in order to subtract baseline drift, non-specific binding and particularly bulk refractive index changes. This reference subtraction method will not be effective if a large dead volume, and hence significant dispersion, exists between both flow channels. For example, if a device has a dead volume of >0.5 μL between each of four flow cell channels where the flow cells themselves have a volume of about 30-60 nL, the dead volume is as much as 16-fold greater than the flow cell volume. This dead volume must be completely washed out before sample dispersion is no longer present. If we assume that a transition to 95% sample is acceptable, then a wash-out volume of 2.7 times the dead volume is required (note: this dilution process follows a natural logarithm, where e=2.72, hence a volume of sample equal to approximately 2.72 times the dead volume must pass through before the sample concentration reaches about 95% of its actual concentration). Therefore, given a 0.5 μL interchannel dead volume between channels 1 and 2, if performing a binding experiment where the sample is injected at 20 μL/min, the first 1.36 μL of sample entering channel 2 will suffer from dispersion, representing four seconds of data where dispersion differences exist between channels 1 and 2. The dispersion period preferably should be reduced to less than 0.5 seconds in order to measure the kinetic parameters of fast interactions. For example, a very weak interaction will often reach steady state in 1 second or less and dissociation may also occur in less then 1 second. In this case, kinetic resolutions would require reducing the dispersion period to a fraction of second or a fraction of the time of the events to be measured. Thus, a transition of 0.1 to 0.2 seconds is preferable.
Prechannel dead volumes that exist in the flow path before reaching the flow cell will cause dispersion that occurs equally in both channels 1 and 2 and is therefore subtracted during reference curve subtraction. Therefore this is of lesser concern, but it is still advantageous to reduce these dead volumes. If these dead volumes are large, then the injected sample will suffer from dispersion for a considerable portion of the injection. Usually injections of >2 minutes are employed. A prechannel dead volume of about 1 μL will show dispersion for 8 seconds at a flow rate of 20 μL/min. That is about 6.6% of the total injection time assuming a 2 minute injection period. During this 8 second dispersion period the sample concentration will be variable and a similar 8 second dispersion period will occur on ending the injection. These dispersion periods complicate kinetic model fitting to the dispersed regions and should be reduced where possible.
U.S. Pat. No. 6,200,814 describes a means of reducing or eliminating the dispersion due to dead volumes that occur before sample injection and also the interchannel dead volume (see, also, PCT publication WO 03/102580 A1). In the '814 patent the sample is allowed to flow into the flow cell but is prevented from contacting the sensing regions due to the presence of a second, or third, stream of liquid (i.e. running buffer) flowing in the same direction adjacent to the sample stream. Here the non-mixing behavior of parallel flowing laminar flow streams is exploited wherein the sample flow stream is separated from the sensing surface along the flow cell by providing a buffer flow along the sensing region such that a buffer-sample interface extends along the length of the flow cell. In this way the dispersed region is allowed to flow to waste before contacting the sensing area and an effective injection point is moved into the flow cell itself. By changing the relative flow rates between these adjacent laminar flow streams, the sample stream is forced to make a rapid lateral displacement allowing a rapid (<0.3 seconds) transition from running buffer to 100% sample over the chosen sensing area. This method is critically dependent on the non-mixing behavior of laminar flow streams. Any disruptions caused by bubbles will prevent normal operation making the system less robust, requiring expensive degassing devices. Furthermore a means of controlling the flow rate in multiple streams adds additional pump drives that introduce considerable expense.